I. Field of the Invention
The present invention relates generally to gamma ray imaging detectors and, more particularly, to such a detector for use in nuclear medicine applications.
II. Description of the Prior Art
In nuclear medicine, a radioactive isotope is introduced into the area of the body under examination. The radioactive isotope generates gamma rays at certain energies dependent upon the particular isotope used. For example, .sup.99m T.sub.c generates gamma rays having an energy of 140 keV which is used in many nuclear medicine applications.
A collimator is then positioned between the object under study and the gamma ray imaging detector. The collimator typically comprises a lead plate having a plurality of parallel throughbores so that the gamma rays which pass through the collimator to the imaging detector are essentially parallel to each other.
A single sodium iodide crystal doped with thallium is contained within the housing for the image detector so that gamma rays passing through the collimator pass through an entrance window on the housing and impinge upon the crystal. A thin aluminum sheet across the entrance window hermetically encloses the crystal to protect it from moisture and the elements while a light reflective surface is positioned between the window and the crystal.
The sodium iodide crystal forms the scintillator material and has a typical thickness of between 0.25 and 0.50 inches. The crystal in turn is glued to a thick glass sheet known as the light pipe. The light pipe not only transmits photons which are generated by the crystal, but also mechanically supports the crystal.
In order to detect the emission of photons from the crystal, an array of photomultiplier tubes (PMT) are optically coupled to the light pipe assembly through an optical coupling compound or cement.
In operation, gamma rays generated by the source pass through the collimator, through the entrance window and are absorbed by the sodium iodide crystal. Only a very few gamma rays are absorbed by the aluminum sheet across the entrance window due to the low atomic number of aluminum. Conversely, the relatively high atomic number of the crystal makes it a good absorber of gamma rays.
Most of the gamma rays that are of interest to nuclear medicine fall within the energy range of 60-360 keV. These gamma rays interact with the crystal through the photoelectric effect in which a bound electron in the crystal absorbs the gamma ray. Upon doing so, the entire energy of the gamma ray is transferred into kinetic energy of the electron.
The electron which absorbs the energy of the incoming gamma ray transfers this energy to adjacent electrons through coulomb collisions. Many electron-hole pairs are formed in the crystal as a result of the energy deposited by each gamma ray photon. The recombination of the holes in the electrons then creates a large number of low energy scintillation photons. Measurements have shown that for a sodium iodide scintillation crystal, 11.4-13.5% of the total energy from the absorbed energy is emitted as scintillation photons.
The scintillation photons are emitted in random directions by the crystal. Those photons striking the light reflective surface are reflected back towards the photomultiplier tubes so that for each gamma ray absorption, photons are emitted in a conical pattern towards the photomultiplier tubes with a higher concentration of the photons in the center of the cone. These photons are detected by the photomultiplier tubes and, for best position resolution of the gamma ray absorption, the photons should strike at least seven photomultiplier tubes. Well known electronic circuitry is then employed to determine the position of the gamma ray absorption from the output signals of the photomultiplier tubes.
These previously known gamma ray imaging detectors, however, suffer from a number of disadvantages. One disadvantage is that the sodium iodide crystal generates only a relatively small number of scintillation photons per absorbed gamma ray. For example, assume that the incoming gamma ray has an energy of 140 keV which is the energy of the .sup.99m T.sub.c decay commonly used in nuclear medicine applications. Absorption of such a gamma ray by the sodium iodide crystal would generate about 5,300 photons with a standard deviation according to Poisson statistics of 73 photons. Thus, for such a gamma ray, 5,300.+-.73 total photons are emitted per gamma ray absorption in a spectrum centered about 415 nanometers which is in the bluish-green visible range.
The scintillation photons travel radially outward from the point of absorption in random directions. The photons travelling towards the photomultiplier tubes are refracted at the crystal/light pipe interface due to a mismatch in the refractive indices for sodium iodide (N=1.85) and glass (N=1.5). As previously described, still other photons reflected from the light reflective surface on the anterior side of the crystal so that the photons form a cone of light exiting the crystal/light pipe assembly and having a diameter larger than the diameter of a single photomultiplier tube. Furthermore, for best results, the cone of light exiting the light pipe assembly produces signals in at least seven photomultiplier tubes, i.e. the center and six nearest photomultiplier tubes. In the well known fashion, the magnitude of the signal from each photomultiplier tube is proportional to the number of photoelectrons produced in each photomultiplier tube. These seven signals are then combined to provide a two dimensional position signal indicative of the point of the gamma ray absorption. Digital corrections are then made to the position signal to correct non-linearities and component variations.
Not all of the photons which strike the photomultiplier tubes generate a photoelectron from the photocathode. Instead, the quantum efficiency of the photomultiplier tube is expressed as a percentage, i.e. the percentage of the number of photons striking the photomultiplier tubes which generate a photoelectron from the photocathode. The quantum efficiency for a photomultiplier tube having a bialkalide photocathode is typically about 15-30% when averaged over the emission spectra of common scintillators.
The two most important intrinsic performance characteristics of the previously known gamma ray imaging detectors are the intrinsic spatial resolution and the intrinsic energy resolution. The previously known gamma ray detectors are capable (when measured using 140 keV gamma rays) of about 4 mm full width at half maximum (FWHM) intrinsic spatial resolution and 9.5-12% FWHM energy resolution for 140 keV gamma rays. Furthermore, these two parameters dominate the performance characteristics of the previously known gamma ray detectors.
For example, the energy resolution of a 140 keV gamma ray absorption by the previously known sodium iodide (Tl) crystal is dominated by Gaussian statistics. Application of the Gaussian statistical model reveals that one standard deviation equals the square root of the mean number of photoelectrons produced per gamma ray absorption. Assuming that the energy resolution is 10% FWHM, then a single standard deviation is equal to 4.25% and the mean number of photoelectrons produced per gamma ray absorption is 552. Consequently, for the 5,300.+-.73 photons emitted per gamma ray absorption, and assuming that the quantum efficiency of the multiplier tube is 20-22%, only about half of the scintillation photons created during the gamma ray absorption actually reach the photocathodes of the photomultiplier tubes.
These previously known gamma cameras suffer from a number of disadvantages. One disadvantage of these previously known cameras is the inability to accurately discriminate between "good events" and "bad events". A good event, of course, is a gamma ray which passes directly from the isotope during radioactive decay through the collimator and to the sodium iodide crystal. "Bad events", on the other hand, are gamma rays which are scattered before reaching the camera as well as other incident gamma rays.
In order for the gamma ray camera to differentiate between "good events" and "bad events", it is necessary to establish an energy window or criteria for the photomultiplier tubes in order to differentiate between good and bad events. Thus, gamma ray absorptions which fall within the energy window are considered to be "good events" while events which fall outside the energy window are considered to be "bad events" and are not processed.
The previously known gamma cameras utilizing photomultiplier tubes and sodium iodide crystals, however, necessitate a relatively wide energy window or energy acceptance criteria. This relatively wide energy window which is necessitated by the 9.5-12% FWHM energy resolution is due in large part to the relatively low photoelectron yield of sodium iodide scintillation light when coupled with the low quantum efficiency of the photomultiplier tubes and the refractive index mismatch between the crystal and the light pipe previously discussed. Consequently, some "bad events" are erroneously accepted as "good events" since the bad event nevertheless falls within the relatively broad energy window or acceptance criteria. All of this detracts from the imaging clarity, accuracy and capability of the gamma camera.
It is not possible or practical with these prior art devices to simply narrow the energy acceptance criteria to eliminate the erroneous acceptance and processing of bad events since, to do so, would undesirably eliminate processing of many "good events" due to the intrinsic energy resolution of the NaI(Tl) crystal.
A still further disadvantage of these previously known gamma cameras is that the sodium iodide crystals must be hermetically sealed by the manufacturer in order to protect the crystal from humidity absorption. The sodium iodide crystals are also brittle and can be easily fractured by temperature or thermal shock.